Calcium hydroxyapatite [HAp, Ca10(PO4)6(OH)2], the main inorganic component of the hard tissues in bone and teeth, is a member of the apatite family. Biological apatite comprises the mineral phase of calcified tissues (enamel, dentin and bone) and is observed to be carbonate substituted, calcium deficient. They differ from pure HAp in stoichiometry, composition, crystallinity and other physical and mechanical properties. For in vivo application, all metallic biomaterials have to be biocompatible. The difference in corrosion resistance, mechanical properties and commercial availability decides their application area. Metallic materials like titanium (Ti) metal and its alloys, cobalt (Co)-chromium (Cr)-molybdenum (Mo) alloy and stainless steel (316 L) are widely used as orthopedic and dental implants. Among these, titanium alloy exhibits superior corrosion resistance and excellent mechanical properties and gives better performance as a metallic stem of hip-joint prosthesis but because of prohibitive cost it has limited use in under developed or developing countries including India. So, this cannot be the material of choice for common people of our country. As an alternative, stainless steel (316L) (composition: 0.03% carbon, 2% manganese, 17-20% chromium, 12-14% nickel, 2-4% molybdenum and small amounts of phosphorous, sulphur and silicon) is widely used hip joint implant.
Insertion of a biomaterial in a living tissue creates an artificial interface between the living tissue and the biomaterial that may render primary or secondary reactions leading to changes in the biological system and the implant material. Under such circumstances, porous hydroxyapatite coatings that are osteoconductive increase the speed and strength of bone attachment compared to uncoated implants of the same design. Other than this, it shields the metallic implant from environmental attack or leaching effects and thus minimizes adverse reactions. Wide spectra of methods have been applied for coating of HAp on metals and other substrates (ceramics, polymers and composites), e.g., dip, plasma spraying, electrophoretic deposition, sputter coating, hot isostatic pressing and ion assisted sputtering etc. Of these, plasma spraying has been used as a major technique in applying hydroxyapatite coatings on metal implants to improve implant fixation and bone growth. Other than being expensive, it suffers from several drawbacks.
Reference may be made to the publication of Serekian P., ‘Hydroxyapatite coatings in orthopedic surgery’, pp. 81-97, edited by Geesink R. G. T and Manley M. T., Raven Press Ltd., New York, 1993, wherein the advantages and drawbacks of plasma and flame spraying, electrophoresis, dip coating, magnetron sputtering have been discussed.
Reference may be also be made to Cheang P. and Khor K. A., ‘Addressing processing problems associated with plasma spraying of hydroxyapatite coatings’, Biomaterials, 17, 537-544, 1996, wherein the problems pertinent to the plasma-sprayed HAp coating has been mentioned to be generation of an amorphous phase along with other non-bioactive calcium phosphate phases. The presence of an amorphous phase in the coating is undesirable because natural bone is crystalline, so integrity of the bone-implant interface is compromised. Mechanical tests show failure of the bone-coating-implant interface occurred due to strong resorption and degradation of the coating with high amorphous phase content. Reference may also be made of Barrere F., Blitterswijk van C. A., Groot de K., Layrolle P., ‘Influence of ionic strength and carbonate on the Ca—P coating formation from SBF×5 solution’, Biomaterials, 23, 1921-1930, 2002 and Liu Y., Layrolle P., Bruijn de J., Blitterswijk van C. and Groot de K., ‘Biomimetic coprecipitation of calcium phosphate and bovine serum albumin (BSA) on titanium alloy’, J. Biomed. Mater. Res., 57[3], 327-335, 2001, wherein drawbacks of plasma-sprayed HAp coatings have been related to extremely high processing temperature (>10,000° C.) that could not produce bone like apatite, coat heat-sensible complex shaped porous implants or incorporate biologically active molecules such as osteogenic agents and growth factors that increase bone regeneration and have osteo-inductive effect.
Recently, an emerging technique, called biomimetic coating overcomes all the intrinsic drawbacks of the plasma-spraying method. It elaborates a dense, uniform and homogeneous hydroxyapatite (bone-like, carbonate substituted, calcium deficient) coating on metal substrates under mild conditions of pH and at room temperature, in simulated body fluid (SBF) that has similar inorganic composition as human blood plasma. Before applying coating, the substrates are preferably cleaned or treated to remove any surface contaminants and to promote good adhesion of the coating. The metallic implants may be rinsed with a degreaser such as acetone, alkyl alcohols etc. followed by using deionised water. To improve coating adhesion, mechanical surface treatments e.g., sand blasting, scouring, polishing and grinding increase surface roughness and improve bond strength between coating and substrate. Chemical surface treatments are also applied with similar purposes. Acid etching using strong mineral acids e.g., hydrofluoric, hydrochloric, sulfuric, nitric, perchloric or oxidising agents like nitric acid, peroxyhalogen acid, hydrogen peroxide etc. form a fresh bioactive metal oxide layer. After the mechanical or chemical treatment, the surface contaminants are removed by rinsing the implant with deionised water under ultrasound.
Reference may be made to U.S. Pat. No. 5,068,122, Kokubo T., Yamamuro T. and Yoshio A., ‘Process for forming a bioactive hydroxyapatite film’, 1991, wherein a process for applying a bioactive hydroxyapatite film on inorganic, metallic or organic implant substrates by soaking an assembly comprising a glass (CaO and SiO2) facing a substrate at a predetermined distance apart in an aqueous solution supersaturated with constituent ions of hydroxyapatite has been discussed. In the method according to the present invention, it is not necessary to provide an assembly of glass facing the substrate to be coated.
Reference may be made to U.S. Pat. No. 6,569,489, Li P., ‘Bioactive ceramic coating and method’, 2003, wherein a bioactive bone mineral (dense, carbonated apatite, crystal size<1 μm) is chemically bonded to a variety of substrates (silicon, metals, ceramics, and polymers), for application in orthopedic and dental prostheses. This coating (thickness 0.005 to 50 μm) is applied uniformly to substrate surfaces of varying geometry and surface textures. It is firmly secured to the substrate and encourages rapid and effective bone ingrowth. The coating is applied by immersing the substrate in an aqueous solution containing calcium, phosphate and carbonate ions (pH range 5-10, temperature<100° C.). Other ions, such as sodium, potassium, magnesium, chloride, sulfate, and silicate, may optionally be present in the solution. The solution is exposed in a controlled environment when it reacts with the substrate to form the coating. The synthetic bone apatite film produced by this process, results in an effective bone composition that promotes bone ingrowth and thereby provides implants with bone-bonding properties. The synthetic apatite film can also be used to attract biological molecules such as growth factors for further improvement of bone growth.
Reference may be made to U.S. Pat. No. 6,733,503, Layrolle P. J. F., de Groot K., de Bruijn J. D., van Blitterswijk C. A., Huipin Y., ‘Method for coating medical implants’, 2004, wherein a bioactive carbonated calcium phosphate layer for coating the surface of medical implants (stainless steel, titanium, nickel, cobalt, chrome, niobium, molybdenum, zirconium, tantalum and their alloys, alumina, zirconia, bioactive glasses, calcium phosphates) has been provided in an improved bioreactor or fermentor system. After cleaning and acid etching, soaking the implantable device into highly concentrated calcifying solution at low temperature produces the coating. Calcium, phosphate, magnesium, carbonate and additionally sodium chloride salts are dissolved in water by bubbling carbon dioxide gas that increases pH and saturation until there is nucleation of the carbonated calcium phosphate crystals on the surface of the implantable device.
Reference may be made to U.S. Pat. No. 6,692,790, Liu Y., Groot de K., Layrolle and P. J. F., ‘Proteinaceous coating’, 2004, wherein an implant material (stainless steel, titanium, nickel, cobalt, chrome, niobium, molybdenum, zirconium, tantalum and their alloys, alumina, zirconia, bioactive glasses, calcium phosphates) is cleaned and surface treated (acid etching) followed by immersion in an aqueous solution comprising a protein, (albumin, caesin, gelatin, lysosime, fibronectin, fibrin and chitosan) calcium and phosphate ions through which a gaseous weak acid is passed, degassing the solution. The coating is precipitated on the implant followed by submersing the coated implant into a second solution to redissolve the magnesium, calcium and phosphate ions and to obtain the proteinaceous coating. Here, uniform precipitation of a calcium phosphate layer on the implant surface are formed under modulated nucleation and crystal growth conditions that mimics the way hydroxyapatite crystals are formed in the body. In this method, considering the physiological conditions under which the biomimetic coating is a grown, biologically active agent such as antibiotics can be coprecipitated.
Reference may be made to Jonasova L., Muller F. A., Helebrant A., Strnad J. and Peter G., ‘Biomimetic apatite formation on chemically treated titanium’, Biomaterials, 25, 1187-1194, 2004, wherein acid etching of the titanium alloy in HCl under inert (CO2) atmosphere for 2 hours was done to obtain a uniform initial substrate surface before alkali treatment. This resulted in formation of a micro-roughened surface and degradation of the passive oxide layer on the titanium alloy (Ti6Al4V) surface,TiO2+4HCl→TiCl4+2H2O  (1)
Next, alkali treatment with 10 M NaOH at 60° C. for 24 hours resulted in dissolution of the passive TiO2 layer and formation of an amorphous layer containing alkali ions,TiO2+NaOH→HTiO3−+Na+  (2)
When soaked in 1×SBF at 37° C. for 20 days, Na+ ions from the amorphous layer were exchanged by H3O+ ions from the surrounding fluid resulting in a Ti—OH surface layer formation. This incorporates the Ca2+ ions that in turn act as nucleation sites for a homogeneous and thick hydroxycarbonated (HCA) apatite layer formation with bone bonding ability by attaching (PO4)3− and (CO3)2− to form Ca—P enriched surface layer. It was seen, thickness of the precipitated HCA layer increased continuously with time.
SBF was prepared by dissolving the reagent grade salts as is mentioned in the above reference (Jonasova L. et al, ‘Biomimetic apatite formation on chemically treated titanium’, Biomaterials, 25, 1187-1194, 2004), in distilled water and buffered at pH 7.3 with tris-hydroxymethyl aminomethane and HCl at 37° C.
Reference may be made to Song W-H., Jun Y-K., Han Y., Hong S-H., ‘Biomimetic apatite coating on micro-arc oxidized titania’, Biomaterials, 25, 3341-3349, 2004, wherein biomimetic apatite coatings on micro-arc oxidized titania films were investigated and their apatite-inducing ability was evaluated in both 1×SBF and 1.5×SBF. 1.5×SBF was prepared using the same reagents as above with ion concentrations 1.5 times 1×SBF. When immersed in 1×SBF carbonated hydroxyapatite was induced on the surfaces of the films oxidized at higher voltages (>450V) after 28 days whereas the use of 1.5×SBF reduced the apatite induction time and apatite formation was confirmed even on the surface of the films oxidized at 350 V.
Reference may also be made to Barrere F., Blitterswijk van C. A., Groot de K. and Layrolle P., ‘Influence of ionic strength and carbonate on the Ca—P coating formation from SBF×5 solution’, Biomaterials, 23, 1921-1930, 2002, wherein biomimetic calcium phosphate coatings were applied on Ti6Al4V alloy using SBF concentrated by a factor 5 (5×SBF) that reduced the soaking time to less than 24 hours. Merck grade reagents were used here and pH was reduced to 6 using CO2 gas at 37° C. Also, it was observed that the coating deposition kinetics is influenced by ionic strength of the solution and HCO3− content. HCO3− reduces the apatite crystal size of the coating allowing better physical attachment on the titanium alloy substrate. Similarly, reference may be made to Barrere F., Valk van der C. M., Meijer G., Dalmeijer R. A. J., Groot de K., Layrolle P., and ‘Osteointegration of biomimetic apatite coating applied onto dense and porous metal implants in femurs of goats’, J. Biomed. Mater. Res. Part B: App. Biomaterials, 67B (1), 655-665, 2003, wherein a 30 μm thick carbonated apatitic coating was developed on porous Ti-6Al-4V and tantalum cylinders by immersion into 5×SBF at 37° C. then at 50° C. for 24 hours. These were implanted in the femoral diaphysis of female goats and bone contact was found higher for the above coated implants in comparison to the uncoated implants.
Reference may be made to Tas A. C. and Bhaduri S. B., ‘Rapid coating of Ti6Al4V at room temperature with a calcium phosphate solution similar to 10× simulated body fluid’, J. Mater. Res., 19[9], 2742-2749, 2004, wherein a 20-65 μm thick bone-like apatitic calcium phosphate coating has been formed on Ti6Al4V substrate at room temperature in 2-6 hours using a super strength solution having concentration of calcium and phosphate ions of human blood plasma/SBF multiplied by a factor of 10. Using a lower concentration (e.g., 1.5×SBF) of SBF, a longer time of two to three weeks is required for the calcium phosphate coating formation in the same method. No buffering agents have been used, instead, prior to coating the pH has been adjusted to 6.5 by addition of sodium bicarbonate. The obtained adhesion strength (12±2 MPa) is comparable to the coating formed by soaking in 1.5×SBF and the Ca/P molar ratio is 1.57.
Reference may also be made to Lin F. H., Hsu Y.-S., Lin S.-H., Sun J.-S., ‘The effect of Ca/P concentration and temperature of simulated body fluid on the growth of hydroxyapatite coating on alkali-treated 316 L stainless steel’, Biomaterials, 23, 4029-4038, 2002, wherein the importance of stainless steel (316L) in orthopedics and dentistry has been discussed. In this, stainless steel (316L) metallic substrates were soaked in 10 M NaOH (aq. soln.) at 60° C. for 24 hours followed by washing with distilled water and drying at 40° C. for 24 hours in air. These were then heated to 600° C. for 1 hour leading to formation of a thin linking layer of sodium chromium oxide between the HAp coating and the metal substrate. This alkali treated substrate was soaked into SBF (equivalent composition of human blood plasma) at a temperature of 80° C. to yield a dense and uniform bone-like HAp layer on the surface in a period of one week. On increase of calcium and phosphorous ions in the coating, iron oxide and iron chromium oxides were formed on the surface that loosens the HAp layer.
Reference may again be made of Teixeira R. L. P., de Godoy G. C. D. and Pereira M. de M., ‘Calcium phosphate formation on alkali-treated titanium alloy and stainless steel’, Materials Research, 7[2], 299-303, 2004, wherein biomimetic method has been adopted for coating (calcium phosphate) alkali-treated titanium and stainless steel alloys as an alternative to plasma spraying method. Here, a comparative study has been pursued on biomimetic coating of HAp on alkali treated AISI 316 L and Ti6Al4V alloy substrates. Best results were obtained in 5N sodium hydroxide (NaOH) salt treated (kept at 60° C. for 24 hours) and 20 N NaOH treated stainless steel (kept at 90° C. for 30 minutes). The titanium samples were heat treated at 600° C. for 1 hour and stainless steel samples at 900° C. for ½ hour in air followed by cooling and soaking in SBF at 37° C. for 3 days and in 1.5 SBF at 37° C. for 1 week. This led to the nucleation of a thin (XRD data) calcium phosphate film on the preformed precursors of sodium chromate layer on stainless steel and sodium titanate layer on titanium alloy. An increase in central roughness was noticed with increase in NaOH concentration. This affected the adhesion of the coating through a micro-mechanical adhesion mechanism, though the previous studies have shown that adherence of biomimetic coating increases with alkali treatment of titanium substrates.
Incorporation of biological moieties in the surface treatment step has osteoinductive effect and increase bone regeneration. By suitable choice of concentration of a water-soluble protein, a stable monolayer (held by secondary forces at the inter- and intramolecular level) is adsorbed slowly (evident by the change in pH of the protein solution) on the mechanically roughened (textured) surface of the metal implant at room temperature. This in turn acts as a functionalized template for the nucleation of the requisite calcium phosphate phase (HAp) as a result of electrostatic interactions. The novelty of the present invention lies in development of a single phase HAp coating on the functionalized biomolecular template of the metal substrate at room temperature (35-37° C.). The coating has increased thickness and increased crystallinity of the as-prepared calcium deficient carbonate substituted hydroxyapatite phase. There are alterations in crystal geometry and indication of crystal growth in a preferred orientation. The coating has uniform coverage, porosity and biocompatibility lowering the surface treatment time from more than forty eight hours as is there in the reported processes to only to four hours. Modified crystallinity and physical structure indicates the effect of the underlying biomolecular template that is interfering with the crystal growth and gets incorporated into the mineral latticework. Reference may be made to Wen H. B., Wijn de J. R., Blitterswijk van C. A. and Groot de K., ‘Incorporation of bovine serum albumin in calcium phosphate coating on titanium’, J. Biomed. Res., 46[2], 245-252, 1999, wherein Ca—P coatings at physiological temperature (37° C.) has been produced on titanium implant materials. This is based on complicated and time consuming wet chemical techniques involving acid etching, boiling diluted alkali incubation followed by immersion in a super saturated calcification solution. Hence, a protein delivery system is produced by coprecipitation of osteogenic proteins (bovine serum albumin) in the coating was chosen in this for coprecipitation and release of biologically active proteins in vivo. Reference may also be made of Liu Y., Layrolle P., Bruijn de J., Blitterswijk van C. and Groot de K., ‘Biomimetic coprecipitation of calcium phosphate and bovine serum albumin on titanium alloy’, J. Biomed. Res., 57[3], 327-335, 2001, wherein biomimetic coprecipitation of bovine serum albumin (BSA) and calcium phosphate on the titanium alloy surface was done or BSA was deposited onto a biomimetically preformed calcium phosphate matrices. In these methods, no reference of the porosity that is necessary for osteointegration was given therein. Also, the calcium phosphate ceramic coating precipitated on the titanium alloy substrate has been used basically as a carrier system for controlled release of biologically active/osteogenic agents in vivo.
The main drawbacks of the above known art and processes are:    1. Poor coating thickness, a maximum of 65 μm could be achieved in comparison to an average thickness of 100-150 μm in case of plasma-sprayed method    2. Multistep time consuming surface treatment process of the metal substrates    3. Details of porosity parameters (pore size, distribution) of the HAp coating could not be obtained that is necessary for osteointegration    4. Formation of amorphous/poorly crystallized carbonated, calcium deficient hydroxyapatite coating.